Radiograms have been produced by using layers of radiation sensitive materials to directly capture radiographic images as image-wise modulated patterns of electrical charges. Depending on the intensity of the incident X-ray radiation, electrical charges generated either electrically or optically by the X-ray radiation within a pixelized area are quantized using a regularly arranged array of discrete solid-state radiation sensors.
U.S. Pat. No. 5,319,206 describes a system employing a layer of photoconductive material to create an image-wise modulated areal distribution of electron-hole pairs which are subsequently converted to corresponding analog pixel (picture element) values by electro-sensitive devices, such as thin-film transistors. U.S. Pat. No. 5,262,649 describes a system employing a layer of phosphor or scintillation material to create an image-wise modulated distribution of photons which are subsequently converted to a corresponding image-wise modulated distribution of electrical charges by photosensitive devices, such as amorphous silicon photodiodes. These solid-state systems have the advantage of being useful for repeated exposures to X-ray radiation without consumption and chemical processing of silver halide films.
In systems utilizing a photoconductive material such as selenium such as the prior art, a conventional radiation imaging system 100 shown in FIG. 1, before exposure to image-wise modulated X-ray radiation, an electrical potential is applied to the top electrode 110 to provide an appropriate electric field. During exposure to X-ray radiation, electron-hole pairs are generated in the photoconductive layer 190, under the dielectric layer 120, in response to the intensity of the image-wise modulated pattern of X-ray radiation, and these electron-hole pairs are separated by the applied biasing electric field supplied by a high voltage power supply. The electron-hole pairs move in opposite directions along the electric field lines toward opposing surfaces of the photoconductive layer 190. After the X-ray radiation exposure, a charge image is received at the charge-collection electrode 130 and stored in the storage capacitor 160 of the transistor 150, which is formed on the substrate 170. This image charge is then readout by an orthogonal array of thin film transistors and the charge integrating amplifier 140. This type of direct conversion system has the distinct advantage of maintaining high spatial resolution more or less independent with the thickness of the x-ray converting photoconductive layer. However, currently, only a very limited number of direct converting photoconductors can be used for commercial products.
The most popular and technical matured material is amorphous selenium that has good charge transport properties for both holes and electrons generated by the x-ray. However, selenium having an atomic number of 34 has only good x-ray absorption in the low energy range, typically below 50 KeV. The absorption coefficient of selenium at higher energy x-ray is smaller and therefore thicker selenium layer is required for adequate x-ray capture. Since the complication and difficulty of fabrication of good imaging quality amorphous selenium is a strong function of the selenium thickness, successful x-ray imaging products are limited to lower energy x-ray applications such as mammography, low energy x-ray crystallography, and low energy non-destructive testing.
For high energy or high intensity x-ray applications, a large number of electron hole-pairs can be generated from each absorbed x-ray photon. When the electrons and holes move along the electric field to the charge collecting electrodes or to the bias electrode, a significant number of electrons and/or holes can be trapped in the selenium layer. These trapped charges will alter the local electric field, and therefore the subsequent charge transport and charge generation efficiency, resulting in a shadow of the previous image superimposed on the subsequent image in a phenomenon known as “ghosting”. Certain image erasing processes are in general required to remove these charges and to restore the selenium layer to uniform charge conversion properties.
After exposure to a first x-ray, selenium experiences charge trapping, and therefore, it suffers from the ghosting effect. Due to these unwanted results, an era se process is needed to reduce the ghosting. K-band radiation from amorphous selenium can also deteriorate image resolution. Consequently, systems utilizing a photoconductive material between a dielectric layer 120 and a charge-collection electrode 130, such as the prior art shown in FIG. 1, are unable to generate high-quality (e.g. high-resolution) images at high energy ranges of x-rays, such as in the 100 keV-MeV range. In fact, such prior art devices are typically only able to generate high-resolution images in a range up to tens of keVs, such as below 50 keV.
It is therefore desirable to design a radiation imaging system without loss of resolution, and with minimized ghosting in high radiation energy or high dose.
During radiation therapy using charged particles, the patient is in a high-background radiation room (there are significant background x-rays and gamma rays). In such an environment, it is desirable to have a detector that has high detection efficiency for charged particles and low detection efficiency for x-rays or gamma rays.
One method of radiation therapy is proton therapy, in which a beam of high energy protons is directed to a patient. One advantage of proton therapy in providing treatment is that protons deposit the majority of their ionization dose at a particular location in the body and then travel no further through the body. This effect results in less damage to tissue surrounding a target. However, since the proton beam does not travel through the body, in proton therapy, the proton cannot be detected after passing through the patient, and it has been difficult to accurately detect the energy of the proton beam.
There is presently a need for doctors to know whether a proton beam is radiated to a desired location for treatment and whether the intensity of the proton beam is at a desired level.
Conventionally, the detection or measurement of the proton beam being used to treat the patient has not been possible. Instead, a separate proton beam (test beam) is irradiated against a detector, and the location and intensity of the beam are detected. A separate proton beam (treatment beam) is irradiated against the patient for treatment.
FIG. 6 provides an example of such a system. As shown in FIG. 6, the conventional proton beam therapy system 600 includes a scintillation panel 601, a camera 602 and a mirror 603 to direct scintillation (photons) from the scintillation panel 601 to the camera.
After a position and intensity are detected by the system 600, the system 600 may be moved and a treatment beam may be irradiated against a patient. In the alternative, the simulated beam may be generated in parallel with a treatment beam. In either case, it is impossible to have real time detection of the position and intensity of the treatment beam, or “inline dosimetry.” Consequently, there may be differences between the position and intensity of the simulated beam and the treatment beam, and treatment effectiveness may be less effective.